The present invention relates generally to the field of medical devices, and more particularly to guiding means such as a guidewire for advancing a catheter within a body lumen to perform a minimally invasive procedure such as percutaneous transluminal coronary angioplasty (PTCA). The present invention further pertains to catheters and sheaths for delivering and deploying an implantable device within a body lumen.
In a typical PTCA procedure a guiding catheter having a preformed distal tip is percutaneously introduced into the cardiovascular system of a patient by means of a conventional Seldinger technique and advanced proximally until the distal tip of the guiding catheter is seated in the ostium of a desired coronary artery. A guidewire is positioned within an inner lumen of a dilatation catheter and then both are advanced through the guiding catheter to the distal end thereof. The guidewire is first advanced out of the distal end of the guiding catheter into the patient's coronary vasculature until the distal end of the guidewire crosses a lesion to be dilated, then the dilatation catheter having an inflatable balloon on the distal portion thereof is advanced into the patient's coronary anatomy over the previously introduced guidewire until the balloon of the dilatation catheter is properly positioned across the lesion. Once in position across the lesion, the balloon is inflated one or more times to a predetermined size with radiopaque fluid to compress the arteriosclerotic plaque of the lesion against the inside of the artery wall and to otherwise expand the inner lumen of the artery. The balloon is then deflated so that blood flow resumes through the dilated artery and the dilatation catheter is removed.
In a conventional stent delivery procedure, a stent is delivered endoluminally on a delivery catheter, then expanded either by an angioplasty balloon or by removing a constraining sheath and permitting the stent to radially expand by its shape memory, superelastic or self-expanding properties. Conventional guidewires for angioplasty and stent-delivery procedures usually comprise an elongated core member with the distal portion of the core member having one or more tapered sections and a flexible body such as a helical coil disposed about the distal portion of the core member. A shapeable member, which may be the distal extremity of the core member or a separate shaping ribbon which is secured to the distal extremity of the core member extends through the flexible body and is secured to a rounded plug at the distal end of the flexible body. Torquing means are provided on the proximal end of the core member to rotate, and thereby steer, the guidewire while it is being advanced through a patient's vascular system.
Further details of guidewires can be found in U.S. Pat. No. 4,516,972 (Samson); U.S. Pat. No. 4,538,622 (Samson, et al.); U.S. Pat. No. 4,554,929 (Samson, et al.); U.S. Pat. No. 4,616,652 (Simpson); U.S. Pat. No. 4,748,986 (Morrison et al.); U.S. Pat. No. 5,135,503 (Abrams); U.S. Pat. No. 5,341,818 (Abrams et al.); and U.S. Pat. No. 5,411,476 (Abrams et al.) each of which is hereby incorporated herein in their entirety by reference thereto.
A major requirement for guidewires and other intraluminal guiding members, whether they be solid wire or tubular members, is that they have sufficient column strength to be pushed through a patient's vascular system or other body lumen without kinking. However, they must also be flexible enough to pass through tortuous passageways without damaging the blood vessel or other body lumen through which they are advanced. Efforts have been made to improve both the strength and flexibility of guidewires in order to make them more suitable for their intended uses, but these two properties tend to be diametrically opposed to one another in that an increase in one usually involves a decrease in the other.
The prior art makes reference to the use of alloys such as NITINOL (nickel-titanium alloy) which have shape memory and/or superelastic or pseudoelastic characteristics in medical devices which are designed to be inserted into a patient's body. The shape memory characteristics allow the prior art devices to be deformed while in the martensite phase to facilitate their insertion into a body lumen or cavity and then be heated within the body to transform the metal to the austenite phase so that the device returns to its remembered shape or to exert a force on whatever prevents the device from returning to its zero strain configuration. Superelastic characteristics on the other hand generally allow the metal to be deformed and restrained in the deformed condition to facilitate the insertion of the medical device containing the metal into a patient's body, with such deformation causing the phase transformation, e.g. austenite to martensite. Once within the body lumen the restraint on the superelastic member can be removed, thereby reducing the stress therein so that the superelastic member can return to its original undeformed shape by the transformation back to the original austenite phase or so that the superelastic member can exert a force on whatever prevents the superelastic member from returning to its zero strain configuration. In other applications, the stress induced austenite to martensite transformation is utilized to minimize trauma while advancing a medical device such as a guidewire within a patient's body lumen.
Shape memory or superelastic alloys generally have at least two phases, a martensite phase, which has a relatively low strength and which is stable at relatively low temperatures and higher strains, and an austenite phase, which has a relatively high strength and which is stable at temperatures higher and strains lower than the martensite phase. Shape memory characteristics are imparted to the alloy by heating the metal at a temperature above body temperature, preferably between about 40° to about 60° C., while the metal is kept in a constrained shape and then cooled to ambient temperature. The cooling of the alloy to ambient temperature causes at least part of the austenite phase to transform to the martensite phase which is more stable at this temperature. The constrained shape of the metal during this heat treatment is the shape programmed when the alloy is reheated to these temperatures causing the transformation of the martensite phase to the austenite phase. The metal in the martensite phase may be plastically deformed to facilitate the entry thereof into a patient's body. The metal will remain in the pre-programmed shape even when cooled to a temperature below the transformation temperature back to the martensite phase, so it must be reformed into a more usable shape, if necessary. Subsequent heating of the deformed martensite phase to a temperature above the martensite to austenite transformation temperature causes the deformed martensite phase to transform to the austenite phase and during this phase transformation the metal reverts back to its remembered shape or to exert a force on whatever prevents the device from returning to its zero strain configuration.
When stress is applied to a specimen of a metal such as NITINOL® exhibiting superelastic characteristics at a temperature at or above which the transformation of martensite phase to the austenite phase is complete, the specimen deforms elastically until it reaches a particular stress level where the alloy then undergoes a stress-induced phase transformation from the austenite phase to the martensite phase. As the phase transformation proceeds, the alloy undergoes significant increases in strain but with little or no corresponding increases in stress. The strain increases while the stress remains essentially constant until the transformation of the austenite phase to the martensite phase is complete. Thereafter, further increase in stress is necessary to cause further deformation. The martensitic metal first yields elastically upon the application of additional stress and then plastically with permanent residual deformation.
If the load on the specimen is removed before any permanent deformation has occurred, the martensitic specimen will elastically recover and transform back to the austenite phase. The reduction in stress first causes a decrease in strain. As stress reduction reaches the level at which the martensite phase transforms back into the austenite phase, the stress level in the specimen will remain essentially constant (but substantially less than the constant stress level at which the austenite transforms to the martensite) until the transformation back to the austenite phase is complete, i.e., there is significant recovery in strain with only negligible corresponding stress reduction. After the transformation back to austenite is complete, further stress reduction results in elastic strain reduction. This ability to incur significant strain at relatively constant stress upon the application of a load and to recover from the deformation upon the removal of the load is commonly referred to as superelasticity or pseudoelasticity.
The prior art makes reference to the use of metal alloys having superelastic characteristics in medical devices which are intended to be inserted or otherwise used within a patient's body. See for example, U.S. Pat. No. 4,665,906 (Jervis) and U.S. Pat. No. 4,925,445 (Sakamoto et al). The Sakamoto et al. patent discloses the use of a nickel-titanium superelastic alloy in an intravascular guidewire which could be processed to develop relatively high yield strength levels. However, at the relatively high yield stress levels which cause the austenite-to-martensite phase transformation characteristic of the material, it did not have a very extensive stress-induced strain range in which the austenite transforms to martensite at relative constant stress. As a result, frequently as the guidewire was being advanced through a patient's tortuous vascular system, it would be stressed beyond the superelastic region, i.e. develop a permanent set or even kink which can result in tissue damage. This permanent deformation would generally require the removal of the guidewire and the replacement thereof with another. Products of the Jervis patent on the other hand had extensive strain ranges, i.e. 2 to 8% strain, but the relatively constant stress level at which the austenite transformed to martensite was very low, e.g. 50 ksi.
The prior methods of using the shape memory characteristics of these alloys in medical devices intended to be placed within a patient's body presented operational difficulties. For example, with shape memory alloys having a stable martensite temperature below body temperature, it was frequently difficult to maintain the temperature of the medical device containing such an alloy sufficiently below body temperature to prevent the transformation of the martensite phase to the austenite phase when the device was being inserted into a patient's body. With intravascular devices formed of shape memory alloys having martensite-to-austenite transformation temperatures well above body temperature, the devices could be introduced into a patient's body with little or no problem, but they had to be heated to the martensite-to-austenite transformation temperature which was frequently high enough to cause tissue damage and very high levels of pain.
What has been needed and heretofore unavailable is tubular body for intravascular devices, such as guidewires or catheter-sheaths, which have at least a portion thereof exhibiting superelastic and/or shape memory characteristics and which is fabricated by vacuum deposition techniques to provide precise control over the crystalline structure of the material used to fabricate the device.